The phenomea of magnetic resonance has been used for a long time to obtain morphological data of biological specimens. More recently magnetic resonance has been used to obtain interior images of in-vivo biological specimens. The images were obtained after it was discovered that the magnetic gradients could be used to locate the source of the signals in magnetic resonance systems. For example, see the article in Nature (London) 242, 190 (1973) by P. C. Lauterbur.
As is well known magnetic resonance imaging uses a relatively strong static magnetic field having a given direction which is aligned with the Z axis of a cartesian coordinate system. The strong static magnetic field causes the nuclei of certain elements such as hydrogen to align with the field. Subsequently radio frequency pulses of sufficient amplitude and/or time duration are applied to nutate the aligned nuclei. The nutated nuclei precess about the Z axis. the rotational frequency of the Rf precession and the frequency of the Rf pulse is the Larmor frequency. The Larmor frequency is related to the strength of the static magnetic field and of the particular element whose nuclei are being nutated according to the equation:
where: EQU f=.gamma.Bo/2.pi.
.gamma. is a constant depending on the element PA1 B.sub.o is the strength of the magnetic field at the point of operation of the radio frequency pulse, and PA1 .pi. is the constant 3.1416+
After the termination of the Rf pulse the nutated nuclei or spins tend to realign with the static magnetic field. Their movement toward realignment in the magnetic field generates Rf signals also having a Larmor frequency. These signals are known as free induction decay (FID) signals. It is these signals that are received to provide information on the proton density of the element that has been nutated by the Rf pulse.
There are many different methods used for obtaining the FID signals. Among the methods and probably one of the most popular methods at the present time is the spin echo method. These methods are all well known and will not elaborated on herein.
A problem occuring in all known methods of obtaining the effective proton densities with MR acquisition systems is that the intensities obtained are relative, i.e. there are not absolute image intensity figures obtained comparable to "CT" numbers in computerized tomography, for example. That is a unit nuclear magnetic resonance (NMR) signal broadcast by different objects will be detected with different intensities by the same NMR spectrometer. The source of these differences is the proximity of the NMR antenna to the source of the signal. This has the result that the antenna is affected differently (in terms of things such as dielectric coupling, capacitance, etc.) by different objects. Thus, although the same unit signal may be broadcast from the body of a person and from a doped water phantom, the antenna will respond differently to the two signals. In particular, the magnitude of the signals detected by the antenna will be different although the signals were the same when produced.
Magnetic resonance images display the effective proton density of a slice of the sample (patient). This effective proton density depends on two groups of parameters. The first group of parameters are subject dependent and includes such things as the actual density of the spins in the slice, the values of various other physical characteristics of the patient such as spin-lattice relaxation time T1, spin-spin relaxation time T2 and the decay time of the FID signal T*2 The second group of parameters are system dependent and includes such things as the gain of the receiving antenna system, the details of the reconstruction algorithm and the scale of the display system.
While the parameters of the first group have physical significance, the parameters of the second group have a large degree of arbitrariness. Since the parameters of the second group particularly the gain of the antenna (i.e. its response to a unit input signal) often change from scan-to-scan, the image intensity values change arbitrarily from scan to scan.
It is an object of the present invention to provide means for removing the dependence of the intensity on this second group of parameters and thus removing the arbitrariness from the measured image intensity value to thereby provide absolute image intensity values. Thus broadly speaking it is an object of the invention to assure that the measured intensity values are independent of the system parameters.
MRI system characteristics are such that each subject causes a different amount of loading on the Rf coils used for transmitting and receiving the Rf signals. Thus, the Rf signals that are received are not absolute signals but are relative signals. For example, more obese persons have a different loading factor than less obese persons.
There is no known method at the present time of co-relating the data obtained from the more obese person with that of the less obese person.
It has been suggested to use a water phantom in the magnet bore with the patient in order to set the gain of the system to obtain signals from the water for use in deriving absolute signal intensities. However, using such a water phantom in the bore with the patient raises the problem of how to properly locate the phantom. For example, if it is positioned below the table, supporting the patient, then it is in a position where the magnetic field is relatively inhomogenous. It is difficult to position the phantom within the center portion of the body coil which magnetically is the ideal position because that is where the patient is positioned. Thus, those skilled in the magnetic resonance imaging art are still seeking systems for calibrating the acquired data in order to standardize the data in a manner similiar to the standardization of CT data. Ideally such a system should not require any extra scan time.